The invention relates to intravascular blood pumps, and specifically to such a pump having multiple, articulated pumping sections.
A common cause of death or disability from heart disease is inadequate supply of blood from an infarcted ventricle, such as caused by cardiogenic shock. If the diseased heart cannot produce a supply of blood sufficient to keep the patient alive, some form of circulatory assistance is required. Circulatory assistance can also be required in allograft cardiac replacement or heart transplants. Patients can die while waiting for transplants or while immunosuppressive agents combat the body's rejection of the transplanted heart.
Several ventricular assist devices ("VADs") have been proposed. Such devices are reviewed in the journal article Ott, Mills, Eugene, and Gazzaniga, Clinical Choices for Circulatory Assist Devices, XXXVI Transactions, American Society of Artificial Internal Organs (October/December 1990). All of the devices described in this article require major surgery for connection to the vascular system, which can involve unacceptable delay or risk for the patient.
A more acceptable approach that has been proposed, and has been used to a limited extent, is to use an intravascular blood pump. As shown schematically in FIGS. 1 and 2, such a pump 30 can be inserted into the heart 10 through the iliofemoral artery 11, aorta 12, and aortic arch 13. The pump 30 may be disposed in the aortic arch 13 or may be inserted through the aortic semilunar valve 14 into the left ventricle 15. Alternatively, the pump may be inserted into the fight ventricle 15 via the pulmonary artery 17. The pump is driven via a flexible cable 31 from an external power source 32. Blood is drawn from the left ventricle 15 into the pump in inlet flow 34 and is discharged into the aorta in discharge flow 35.
Various intravascular pumps are disclosed in the following U.S. patents:
U.S. Pat. No. 4,625,712 to Wampler (assigned to Nimbus, Inc.) ("Wampler I"); PA1 U.S. Pat. No. 4,817,586 to Wampler (assigned to Nimbus Medical, Inc.) ("Wampler II"); PA1 U.S. Pat. No. 4,846,152 to Wampler, et al. (assigned to Nimbus Medical. Inc.) ("Wampler III"); PA1 U.S. Pat. No. 4,944,722 to Carriker, et al. (assigned to Nimbus Medical, Inc.) ("Carriker"); PA1 U.S. Pat. No. 4,964,864 to Summers, et al. (assigned to American Biomed, Inc.) ("Summers"); PA1 U.S. Pat. No. 4,969,865 to Hwang, et al. (assigned to American Biomed, Inc.) ("Hwang"); and PA1 U.S. Pat. No. 5,040,944 to Cook ("Cook").
Wampler I discloses an axial-flow pump with multiple rows or rotor and stator blades. The rotor rotates at speeds in the range of 10,000 to 20,000 rpm, producing blood flows on the order of 4 liters per minute (l/min). (A healthy heart pumps blood at a rate of between 5 and 9 l/min.) It was recognized in Wampler I that pumps with such high rotational speeds have been considered unsuitable for pumping blood because the shear forces imposed on the blood in the region between the tips of the rotor blades and the stationary wall of the pump chamber can cause severe hemolysis. The solution proposed in Wampler I was to reduce the diameter of the pump chamber to thereby reduce the rotor blade tangential tip velocity and to provide a large clearance between the rotor blade tip and the pump chamber wall. The combination of these two dimensional limitations was considered to produce acceptable shear forces.
However, in Wampler III, these dimensional limitations were considered unsatisfactory. Wampler III identified the design criteria for an intravascular pump to include: a) small diameter (to readily fit into the blood vessels through which it is inserted into the heart); b) short length (to be able to follow sharp bends in the vessels); and c) a minimum number of hemolysis-producing interfaces (regions of high shear forces between rotating and stationary parts). The solution proposed in Wampler III was to use a single-stage pump, which is inherently shorter and has fewer hemolysis-producing interfaces than the multi-stage pump of Wampler I. To obtain the required pressure rise across the single stage without causing stalls or cavitation, the rotor is formed with one row producing flow that is partially axial and partially centrifugal, while the other row produces purely axial flow.
Wampler II proposes a pump design in which the diameter of the pump can be reduced without reducing the diameter of the rotor beating below the practical limits of miniaturization. The solution proposed in Wampler II is to use a screw-type axial-flow pump with a multiple-thread rotor in combination with a housing having radially-directed exit flow slots. The screw-type pump has a cylindrical pump chamber in which a screw-thread rotor rotates, rather than a multi-bladed stator. By discharging the pumped blood radially through the wall of the pump housing, rather than axially through the end of the housing, the rotor beatings can have the same diameter as the housing.
In further pursuit of a shorter pump body, with the stated goal of facilitating the passage of the pump through the bends in the insertion catheter, Carriker proposed a screw-thread type pump with radial exit slots that uses a resiliently extendable rotor extension to connect the rotor to the drive cable.
Summers identified another problem with intravascular pumps driven at high rotational velocities by a drive cable--friction between the drive cable and the wall of the insertion catheter can create hot spots in the femoral arteries. To reduce the pump operating speed, Summers proposed a pump using the moineau pump principle to deliver large volumes of blood at relatively low pressure and rotational velocity. The pump uses rotary motion to move a seal continuously through a resilient stator. Pumping action is achieved by the rotor being driven eccentrically within the stator to form a series of sealded chambers that progress axially along the pump.
In one embodiment, at the discharge end of the stator, a discharge nozzle directs the pumped blood to the intake end of a venturi tube The intake end of the venturi is also open to a chamber formed within the pump housing that communicates with the ventricle through ports in the housing. As the blood pumped by the rotor passes through the venturi, it creates a low pressure region within the venturi tube that draws blood from the chamber into the venturi to mix with the pumped blood before it is discharged from the pump. This produces a higher pumped blood volume at a lower discharge pressure.
The pump disclosed in Summers was purported to pump 3 to 4 l/min at a speed of 2,500 rpm without risk of hemoloysis because the pump had no propellers or turbine blades to produce shear forces on the blood. However, the progressive-chamber pump design requires a relatively long pump to provide adequate flow rates at the desired rotational velocity. The longer pump is more difficult to pass through bends in the insertion catheter.
Another low-rpm pump design was proposed in Hwang. The pump uses a helical-shaped foil rotating within a cylindrical housing. This pump design is purported to pump 3 to 4 l/min at 6,000 to 10,000 rpm. As with Summers, the Hwang pump is relatively long.
Finally, Cook discloses an axial flow pump in which a convoluted stationary member is mounted in a cylindrical housing. A pair of spiral impeller rods are rotatably disposed between the stationary member and the internal wall of the housing and are driven by a conventional drive cable. Again, the pump disclosed is relatively long--1.25" (30 mm)--with a diameter of 0.25" (6 mm) and operates at high speeds--approximately 20,000 rpm.
Another disadvantage shared by all of these pump designs is that they do not allow the pumped blood to be preferentially directed. As shown in FIG. 1, three major arteries branch from the aortic arch 13: the left subclavian artery, 18, the left common carotid artery 19, and the brachicocephalic artery 20. In addition, the coronary arteries 21 and 22, which supply blood to the heart muscle itself, branch from the aortic arch just downstream of the aortic valve 14. It is desirable to preferentially direct some of the discharge blood flow 35 from the pump into the entrances of these arteries. Such directional treatment maximizes the effectiveness of the energy conveyed to the blood from the pump (i.e., minimal energy is wasted through the use of aortic arch walls to redirect discharged blood). As such, the overall efficiency of the pump is enhanced. Further, directional treatment into multiple arteries provides a more uniform and equitable distribution of blood leaving the heart. However, the pump designs described above simply discharge blood along or about the axis of the aorta, with no preferential orientation.